Biosensor with peroxidase enzyme

ABSTRACT

An improved biosensor having at least a first working electrode and a first electrode material disposed on the first working electrode. The first electrode material is a mixture made by combining at least one enzyme where the at least one enzyme is a capable of reacting with the analyte to be measured, a redox mediator capable of reacting with a product of an enzymatic reaction or a series of enzymatic reactions involving the at least one enzyme, a peroxidase capable of catalyzing a reaction involving the redox mediator where the redox mediator is oxidized, a binder and a surfactant.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates generally to a biosensor that can be usedfor the quantification of a specific component or analyte in a liquidsample. Particularly, this invention relates to a new and improvedbiosensor and to a new and improved method of fabricating a biosensorfor the quantification of a specific component or analyte in a liquidsample such as creatinine, creatine, glucose, cholesterol, urea and thelike. More particularly, this invention relates to a disposablebiosensor that is inexpensive to manufacture. Even more particularly,this invention relates to a disposable biosensor and method thataccurately measures various analytes such as creatinine, creatine,glucose, cholesterol and the like in small volume biological fluidsamples. Still even more particularly, this invention relates to amethod of measuring the concentration of various analytes in smallvolume biological fluid samples using a redox mediator and at least anenzyme based on the electrochemical mechanism.

2. Description of the Prior Art

Biosensors have been used in the determination of concentrations ofvarious analytes in fluids for more than three decades. Of particularinterest is the measurement of blood glucose, creatinine, creatine, andcholesterol.

It is well known that the concentration of blood glucose is extremelyimportant for maintaining homeostasis. Products that measurefluctuations in a person's blood sugar, or glucose levels, have becomeeveryday necessities for many of the nation's millions of diabetics.Because this disorder can cause dangerous anomalies in blood chemistryand is believed to be a contributor to vision loss and kidney failure,most diabetics need to test themselves periodically and adjust theirglucose level accordingly, usually with insulin injections. If theconcentration of blood glucose is below the normal range, patients cansuffer from unconsciousness and lowered blood pressure that may evenresult in death. If the blood glucose concentration is higher than thenormal range, the excess blood glucose can result in synthesis of fattyacids and cholesterol, and in diabetics, coma. Thus, the measurement ofblood glucose levels has become a daily necessity for diabeticindividuals who control their level of blood glucose by insulin therapy.

Patients who are insulin dependent are instructed by doctors to checktheir blood-sugar levels as often as four times a day. To accommodate anormal life style to the need of frequent monitoring of glucose levels,home blood glucose testing was made available with the development ofreagent strips for whole blood testing.

One type of blood glucose biosensors is an enzyme electrode combinedwith a mediator compound that shuttles electrons between the enzyme andthe electrode resulting in a measurable current signal when glucose ispresent. The most commonly used mediators are potassium ferricyanide,ferrocene and its derivatives, as well as other metal-complexes. Manysensors based on this type of electrode have been disclosed. Examples ofthis type of device are disclosed in the following patents.

U.S. Pat. No. 5,628,890 (1997, Carter et al.) discloses an electrodestrip having an electrode support, a reference or counter electrodedisposed on the support, a working electrode spaced from the referenceor counter electrode on the support, a covering layer defining anenclosed space over the reference and working electrodes and having anaperture for receiving a sample into the enclosed space, and a pluralityof mesh layers interposed in the enclosed space between the coveringlayer and the support. The covering layer has a sample applicationaperture spaced from the electrodes. The working electrode includes anenzyme capable of catalyzing a reaction involving a substrate for theenzyme and a mediator capable of transferring electrons between theenzyme-catalyzed reaction and the working electrode.

U.S. Pat. No. 5,708,247 (1998, McAleer et al.) discloses a disposableglucose test strip having a substrate, a reference electrode, a workingelectrode, and a means for making an electrical connection. The workingelectrode has a conductive base layer and a coating layer disposed overthe conductive base layer. The coating layer is a filler having bothhydrophobic and hydrophilic surface regions that form a network, anenzyme and a mediator.

U.S. Pat. No. 5,682,884 (1997, Hill et al.) discloses a strip electrodewith screen printing. The strip has an elongated support that includes afirst and second conductor each extending along the support. An activeelectrode, positioned to contact the liquid mixture and the firstconductor, has a deposit of an enzyme capable of catalyzing a reactionand an electron mediator. A reference electrode is positioned to contactthe mixture and the second conductor.

U.S. Pat. No. 5,762,770 (1998, Pritchard et al.) discloses anelectrochemical biosensor test strip that has a minimum volume bloodsample requirement of about 9 microliters. The test strip has a workingand counter electrodes that are substantially the same size and made ofthe same electrically conducting material placed on a first insulatingsubstrate. Overlaying the electrodes is a second insulating substratethat includes a cutout portion that forms a reagent well. The cutoutportion exposes a smaller area of the counter electrode than the workingelectrode. A reagent for analysis of an analyte substantially covers theexposed areas of the working and counter electrodes in the reagent well.Overlaying the reagent well and affixed to the second insulatingsubstrate is a spreading mesh that is impregnated with a surfactant.

U.S. Pat. No. 5,755,953 (1998, Henning et al.) discloses areduced-interference biosensor. The device generally comprises anelectrode used to electrochemically measure the concentration of ananalyte of interest in a solution. The device Includes a peroxidaseenzyme covalently bound to microparticle carbon and retained in a matrixIn intimate contact with the electrode. According to this disclosure, itis the enzyme/microparticle carbon of the device that provides acomposition that displays little sensitivity to known interferingsubstances.

It is well known that creatinine is a waste product derived fromcreatine and excreted by the kidneys. The analytical determination ofcreatinine in urine, serum or plasma is a widely used and extremelyImportant test for renal dysfunction. Measurements of creatinine inserum or urine may also be used as indices in the diagnosis andtreatment of other disorders such as muscular dystrophy andhypothyroidism. Thus, the creatinine assay has been a widely recognizedas having vital medical significance. Further, dietary changes havelittle if any influence on the creatinine concentration in blood andurine. Although creatinine is primarily a waste product, and as suchplays no important role in biochemical functions of the body, itschemical precursor, creatine, has a vital biochemical role. Creatine isa basic building block of creatine phosphate, which is the primary formof energy storage in muscle. As a result, the creatinine level is animportant diagnostic index for renal, muscular and thyroid function.

Spectrophotometry has been conventionally employed for measuringcreatinine. The presence and concentration of creatinine in theabove-mentioned body fluids is most frequently determined by the Jaffereaction. In this reaction, creatinine reacts with picric acid toproduce a red color, a red tautomer of creatinine picrate. This methodsuffers from serious disadvantages including, but not limited to, theinstability of alkaline picrate solutions and the concomitant necessityfor preparing solutions as needed, interference from blood metabolites,the analytical time required to perform the method, and the lack ofspecificity.

Sensors have been developed for the detection of creatinine based onenzymatic cleavage of creatinine. Among them, electrochemical methodsreceived particular attention. Rechnitz et. al. (T. Huvin and G. A.Rechnitz, Anal. Chem., 46 (1974) 246) used creatinine deiminase coupledwith an ammonia electrode to measure ammonia produced by an enzymaticreaction. However, this potentiometric method seems of little usefulnessdue to serious interference problems and the sensitivity limitation ofthe gas-sensing electrode.

U.S. Pat. No. 5,958,786 (1999, C. Munkholm) provides for the coupling ofthe enzymatic cleavage of creatinine to detection by a fluorescentpolymer coating. The polymer coating has a first layer of protonated pHsensitive fluorophore immobilized in a hydrophobic polymer. Thefluorophore reacts quantitatively with ammonia. The transducing moietyof the fluorophore is neutrally charged when deprotonated. The polymercoating has a second layer of creatinine deiminase and a polymer, and athird layer of a polymer. A disadvantage of this device is that twoconsecutive readings must be made. First, a fluorescence measurementmust be made of the creatinine sensor. Second, the sensor material ofthe creatinine sensor is then exposed to a solution containingcreatinine followed by measuring the fluorescence change and determiningthe concentration of creatinine.

A more practical strategy was reported by Tsuchida and Yoda in 1983 (T.Tsuchida and K. Yoda, Clin. Chem., 2911 (1983) 51). The proposed systemconsisted of three enzymes, creatinine amidohydrolase (C1), creatineamidinohydrolase (C2) and sarcosine oxidase (SO). These enzymes wereco-immobilized onto the porous side of a cellulose membrane. Themembrane was combined with a polarographic electrode for sensinghydrogen peroxide, a product resulting from the enzymatic reaction.Several research groups attempted to improve electrode performancethrough better enzyme immobilization techniques. (H. Yamato, M. Ohwa andW. Wemet, Anal. Chem., 67 (1995) 2776; M. B. Madaras, I. C. Popescu, S.Ufer and R. P. Buck, Anal. Chim. Acta, 319 (1996) 335; J. Schneider, B.Grundig, R. Renneberg, K. Camman, M. B. Madaras, R. P. Buck and K. D.Vorlop, Anal. Chim. Acta, 325 (1996) 161). Despite the improvements inenzyme immobilization, the methods suffer from various shortcomingsincluding long-term stability, appropriate dynamic measurement range andserious Interference from other oxidizable substances in the samplefluid such as ascorbic acid and acetaminophen as well as creatine.

Currently, two commercial products for measuring blood creatinine areavailable. One is from Nova Biomedical Corporation. It is a criticalcare analyzer that provides a complete 14-test profile from as little as105 microliters of whole blood where one of the tests is for creatinine.The creatinine sensor is a multiple-use, membrane-based sensor arrangedin a fluid channel along with other biosensors (Nova Stat Profile® M,Nova Biomedical Corporation, Waltham, Mass.). The enzymes areimmobilized onto the membrane and the membrane is attached to theworking electrode (platinum) and the reference electrode (Ag—AgCl).

The second commercial product is from i-Stat Corporation (Kanata,Ontario, Canada). A US patent covers this product. U.S. Pat. No.5,554,339 (1996, Cozzette et al.) discloses an amperometric base sensorfabricated on a planar silicon substrate by means of photolithography incombination with the plasma deposition of metallic substances. Themetallic substances include iridium metal (used as working electrode)and silver metal (served as reference electrode along with resultingchloridized silver). Three enzymes (C1, C2 and SO) are immobilized ontothe electrodes as an overlaid structure. The above two products requirecalibration before measurement and a relatively large amount of samplevolume. They also require a relatively longer waiting time for testresults.

Because of the significance of obtaining accurate analyte concentrationmeasurements, it is highly desirable to develop a reliable,user-friendly and disposable sensor, which does not have all of thedrawbacks previously mentioned. Therefore, what is needed is anelectrochemical sensor that does not require routine maintenance. Whatis further needed is an improved electrochemical sensor that combinesperoxidase with a mediator. What is still further needed is an improvedelectrochemical sensor that combines peroxidase with a mediator and thatoperates at a reductive potential where interferents are not oxidized.What is yet further needed is an improved creatinine electrochemicalsensor that includes an interference-correcting electrode to minimizethe interference effects caused by the presence of creatine in a samplefluid. What is yet further needed are improved electrochemical sensorsfor cholesterol, glucose and other biologically important metabolites.Yet, what is still further needed is an electrochemical sensor thatrequires less sample volume for measuring an analyte than previouslyrequired by the prior art. What is still further needed is an improveddisposable sensor for self-testing.

SUMMARY OF THE INVENTION

It is an object of the present invention to provide an electrochemicalsensor that does not require routine maintenance. It is a further objectof the present invention to provide an electrochemical sensor thatcombines at least one enzyme with a peroxidase and a mediator. It isstill a further object of the present invention to provide anelectrochemical sensor that combines at least one enzyme with aperoxidase and a mediator and that operates at a lower potential whereinterferents are not oxidized. It is yet a further object of the presentinvention to provide a creatinine electrochemical sensor that includesan interference-correcting electrode to minimize the interferenceeffects caused by the presence of creatine in a sample fluid. It is yetfurther object of the present invention to provide improvedelectrochemical sensors for cholesterol, glucose and other biologicallyimportant metabolites. It is yet another object of the present Inventionto provide an electrochemical sensor with high sensitivity to theanalytes to be measured. It is yet still a further object of the presentinvention to provide an electrochemical sensor that requires less samplevolume for measuring analytes than previously required by the prior art.It is still a further object of the present invention to provide animproved disposable sensor for self-testing.

The present invention achieves these and other objectives by providing asimple and convenient method of measuring various analytes in biologicalfluids. Although the following describes a preferred design of thepresent invention, a sensor of the present invention may have differentphysical shapes without detracting from the unique characteristics ofthe present invention. The present invention has a laminated, elongatedbody having a sample fluid channel connected between an opening on oneend of the laminated body and a vent hole spaced from the opening.Within the fluid channel lies one or more working electrodes and areference electrode, depending on the analyte to be measured. Thearrangement of the one or more working electrodes and the referenceelectrode is not important for purposes of the results obtained from thesensor. The working electrodes and the reference electrode are each inelectrical contact with separate conductive conduits, respectively. Theseparate conductive conduits terminate and are exposed for making anelectrical connection to a reading device on the end opposite the openchannel end of the laminated body.

The laminated body has a base insulating layer made from a plasticmaterial. Several conductive conduits are delineated on the baseinsulating layer. The conductive conduits may be deposited on theinsulating layer by screen printing, by vapor deposition, or by anymethod that provides for a conductive layer that adheres to the baseinsulating layer. The conductive conduits may be individually disposedon the insulating layer, or a conductive layer may be disposed on theinsulating layer followed by etching/scribing the required number ofconductive conduits. The etching process may be accomplished chemically,by mechanically scribing lines in the conductve layer, by using a laserto scribe the conductive layer into separate conductive conduits, or byany means that will cause a break between and among the separateconductive conduits required by the present invention. Conductivecoatings or layers that may be used are coatings of copper, gold, tinoxide/gold, palladium, other noble metals or their oxides,or carbon filmcompositions. The preferred conductive coatings are gold film or a tinoxide/gold film composition.

It should be pointed out that although the same electrically conductingsubstance (gold film or tin oxide/gold film) after scoring is used asconducting material for both the one or more working electrodes and thereference electrode, this material itself cannot function as a referenceelectrode. To make the reference electrode work, there must be a redoxreaction (e.g., Fe(CN)₆ ³⁻+e⁻⇄Fe(CN)₆ ⁴⁻ or AgCl+e⁻⇄Ag+Cl⁻) at theelectrically conducting material when a potential is applied. Therefore,a redox reaction must be present at the conducting material used for thereference electrode.

In one embodiment of the present invention, the laminated body has afirst middle insulating layer, also called a reagent holding layer, ontop of the base insulating layer and the conductive conduits. The firstmiddle layer, or reagent holding layer, contains cutouts for one or moreworking electrodes and a reference electrode. Each cutout corresponds toand exposes a small portion of a single conductive conduit. The cutoutsfor the working electrodes are substantially the same size. The cutoutfor the reference electrode may be the same or different size as thecutouts for the working electrodes. The placement of all of the cutoutsare such that they will all co-exist within the sample fluid channeldescribed above. This first middle insulating layer is also made of aninsulating dielectric material, preferably plastic, and may be made bydie cutting the material mechanically or with a laser and then fasteningthe material to the base layer. An adhesive, such as apressure-sensitive adhesive, may be used to secure the first middleInsulating layer to the base layer. Adhesion may also be accomplished byultrasonically bonding the first middle layer to the base layer. Thefirst middle insulating layer may also be made by screen printing thefirst middle insulating layer over the base layer.

Each cutout contains electrode material. The electrode material has aredox mediator and a peroxidase. The peroxidase may be from any sourcesuch as soybean (soybean peroxidase (SBP)) or horseradish root(horseradish root peroxidase (HRP)). For most analytes such as glucoseand cholesterol, at least one of the cutouts contains the electrodematerial and an analyte-related enzyme forming an enzyme mix capable ofcatalyzing a reaction involving a substrate for the enzyme, e.g. glucoseoxidase (GOD) for glucose. The redox mediator is capable of transferringelectrons between the enzyme-catalyzed reactions and the workingelectrode.

For analytes having a substrate capable of undergoing similar reactionsand causing an interference effect, a multiple enzyme mix may berequired. Creatinine is one such analyte. Both creatinine and creatineexist in the blood. To measure the enzyme creatinine using the presentinvention, at least one “working electrode” cutout contains theelectrode material and two enzymes, e.g. creatine amidinohydrolase (C2)and sarcosine oxidase (SO), capable of catalyzing a reaction involving asubstrate for the enzyme creatine. This measures the creatine level. Asecond cutout contains the electrode material and three enzymes, e.g.creatinine amidohydrolase (C1), creatine amidinohydrolase and sarcosineoxidase, capable of catalyzing a reaction involving a substrate for theenzyme creatinine. The difference in output of the two workingelectrodes represents the concentration of creatinine in the samples.

The enzymatic-reaction sequence for a creatinine sensor is:$\begin{matrix}{{Creatinine} + {H_{2}{O\overset{C\quad 1}{\longrightarrow}{Creatine}}}} & {{Eq}.\quad(1)} \\{{Creatinine} + {H_{2}{O\overset{C\quad 2}{\longrightarrow}{Sarcosine}}} + {Urea}} & {{Eq}.\quad(2)} \\{{Sarcosine} + {H_{2}O} + {O_{2}\overset{SO}{\longrightarrow}{Glycine}} + {HCHO} + {H_{2}O_{2}}} & {{Eq}.\quad(3)}\end{matrix}$

Creatinine measurements in the prior art are based on the amperometricdetection of H₂O₂ resulting from the above enzymatic reaction. Theenzymatic-reaction sequence for a glucose sensor is: $\begin{matrix}{{Glucose} + {H_{2}O} + {{O_{2}\overset{GOD}{\longrightarrow}{Gluconic}}\quad{acid}} + {H_{2}O_{2}}} & {{Eq}.\quad(4)}\end{matrix}$

The present invention increases the sensitivity of the analyticmeasurement by incorporating a mediator and a peroxidase enzyme in theelectrode material. The preferable mediators are redox chemicals eitherin oxidized or reduced form. The mediator used in the present inventionmay be at least one of a variety of chemicals in their reduced form, orvirtually any oxidizable species or electron donors. Examples of useablecompounds are Fe(CN)₆ ³⁻, Fe(CN)₆ ⁴⁻, Fe(phen)₃ ²⁺(phen=1,10-phenanthroline), Fe(bpy)₃ ²⁺ (bpy=2,2′-bipyridine), Co(NH₃)₆²⁺, Co(phen)₃ ²⁺, Co(bpy)₃ ²⁺, Os(bpy)₂Cl⁺, Os(phen)₂Cl⁺ Ru(bpy)₂ ²⁺,Rb(bpy)₂ ²⁺, cobalt phthalocyanine, various ferrocenes, methylene blue,methylene green, 7,7,8,8-tetracyanoquinodimethane (TCNQ),tetrathiafulvalene (TTF), toluidine blue, meldola blue,N-methylphenazine methosulfate, phenyldiamines,3,3′,5,5′-tetramethylbenzidine (TMB), pyrogallol, and benzoquinone (BQ).It is desirable that the mediator is capable of being oxidizedchemically by hydrogen peroxide resulting from the enzymatic reactionssuch as those illustrated in Eqs. (1) to (3) and Eq. (4) above. It isfurther desirable that the oxidation form of the mediator is capable ofbeing reduced electrochemically at the working electrodes at the appliedpotential. It is still further desirable that the mediator is stable inthe matrix. The preferred mediator in the present invention is potassiumferrocyanide (K₄Fe(CN)₆).

The reduced form of the ferrocyanide mediator (Fe(CN)₆ ⁴⁻) is capable ofbeing oxidized by the hydrogen peroxide resulting from the aboveenzymatic reaction to Fe(CN)₆ ³⁻ in the presence of a peroxidase. Whenusing ferrocyanide as the mediator, the oxidation reaction is as shownbelow: $\begin{matrix}{{{Fe}({CN})}_{6}^{4 -} + {H_{2}{O_{2}\overset{SBP}{\longrightarrow}{{Fe}({CN})}_{6}^{3 -}}} + {H_{2}O}} & {{Eq}.\quad(5)}\end{matrix}$The oxidized form of the ferrocyanide radical Fe(CN)₆ ³⁻ is capable ofbeing reduced electrochemically when a low potential is applied to theworking electrodes. The resulting current signal is related to theanalyte concentration.

It is well known that dissolved oxygen could be reduced at the electrodewhen a low potential is applied. Thus, it is desirable to apply apotential between the working electrodes and the reference electrodesuch that (Fe(CN)₆ ³⁻ is electro-reduced but dissolved oxygen is not orminimized. Furthermore, it is also desirable to use a potential wherethe electro-oxidation of other oxidizable interferents like ascorbicacid and acetaminophen either does not occur or is minimal. An exampleof such an applied potential is between about 0.0 V and about −0.6 V asmeasured against the reference electrode of the present invention. Thepreferred potential is about −0.15 V. This potential is preferred forproviding a good ratio of signal vs. background noise/interference.

It is also desirable to minimize the interference from hematocrit(volume fraction of erythrocytes) on the results. Because theconductivity (or impedance) of whole blood is dependent on hematocrit,it can then be used to correct the effect of hematocrit on the reportedconcentration.

The resistance (r-value) between W (working electrode) and R (referenceelectrode) is related to the hematocrit as represented by the followingequation:r=k₁/(1−H)  Eq. (6)where r is resistance value measured in Ohms or Kilo-Ohms

-   -   H is hematocrit level    -   k₁ is a constant (r measured in Kilo-Ohms)

The measured “r” can then be used to correct the analyte concentration.The relationship is represented by the Equation (7).:C_(corr) =k₂×C_(mea)×r/r₀  Eq. (7)

-   -   where C_(corr) is the corrected analyte concentration    -   C_(mea) is the measured analyte concentration    -   r₀ is the resistance value in Ohms or Kilo-Ohms measured at a        preselected normal hematocrit    -   k₂ is a constant

The laminated body also has a second middle insulating layer, alsocalled a channel-forming layer, on top of the first middle layer. Thesecond middle layer, or channel-forming layer is also made of a plasticinsulating material and creates the sample fluid channel of thelaminated body. It contains a U-shaped cutout on one end which overlaysthe cutouts on the first middle layer with the open end corresponding tothe open end of the laminated body described earlier.

The laminated body of the present invention has a top layer with a ventopening. The vent opening is located such that at least a portion of thevent opening overlays the bottom of the U-shaped cutout of the secondmiddle insulating layer. The vent allows air within the sample fluidchannel to escape as the sample fluid enters the open end of thelaminated body. The sample fluid generally fills the sample fluidchannel by capillary action. In small volume situations, the extent ofcapillary action is dependent on the hydrophobic/hydrophilic nature ofthe surfaces in contact with the fluid undergoing capillary action. Thisis also known as the wetability of the material. Capillary forces areenhanced by either using a hydrophilic insulating material to form thetop layer, or by coating at least a portion of one side of a hydrophobicinsulating material with a hydrophilic substance in the area of the toplayer that faces the sample fluid channel between the open end of thelaminated body and the vent opening of the top layer. It should beunderstood that an entire side of the top layer may be coated with thehydrophilic substance and then bonded to the second middle layer.

The insulating layers of the laminated body may be made from anydielectric material. The preferred material is a plastic material.Examples of acceptable compositions for use as the dielectric materialare polyvinyl chloride, polycarbonate, polysulfone, nylon, polyurethane,cellulose nitrate, cellulose propionate, cellulose acetate, celluloseacetate butyrate, polyester, acrylic, and polystyrene.

In a second embodiment of the present invention, a first middle layer isnot required for those analyte-measuring electrode systems where thereare no competing substrate reactions for the enzyme. In other words,where there is no need for a second working electrode such as in thecreatinine measuring system of the present invention.

In the embodiments using a first insulating layer, two cutouts containmaterial for the working electrodes (W1 and W2) and one for thereference electrode (R). The positional arrangement of the two workingelectrodes and the reference electrode in the channel are not criticalfor obtaining useable results from the electrochemical sensor. Thepossible electrode arrangements within the sample fluid channel may beW1-W2-R, W1-R-W2, R-W1-W2, W2-W1-R, W2-R-W1, or R-W2-W1 with thearrangement listed as the arrangement of electrodes would appear fromthe open end of the laminated body to the vent opening. The preferredposition was found to be W1-R-W2; that is, as the sample fluid enteredthe open end of the laminated body, the fluid would cover W1 first, thenR, then W2. The working electrodes and the reference electrode are eachin electric contact with separate conductive conduits, respectively. Theseparate conductive conduits terminate and are exposed for making anelectric connection to a reading device on the end opposite the openchannel end of the laminated body.

In the creatinine sensor, the first working electrode (W1) is loadedwith a mixture of C2, SO, a peroxidase, potassium ferrocyanide, at leastone binder, and a surfactant. The second working electrode (W2) isloaded with the same chemical reagent as W1 but with the addition of C1.The reference electrode (R) cutout is loaded with a mixture containingat least one of the redox mediators mentioned above, at least onebinder, and a surfactant. It should be noted that W1 is substantially acreatine sensor, while W2 is substantially a sensor responding tocreatinine and to creatine. The difference between the electroderesponses at W2 and W1 corresponds to the creatinine concentration.

In a glucose sensor, the first working electrode is loaded with amixture of glucose oxidase, a peroxidase, potassium ferrocyanide, atleast one binder, and a surfactant. In a cholesterol sensor, the firstworking electrode is loaded with a mixture of cholesterol esterase,cholesterol oxidase, a peroxidase, potassium ferrocyanide, at least onebinder, and a surfactant. The reference electrode may be loaded with thesame mixture as the working electrode. It should be pointed out that thereference electrode cutout could be loaded with a Ag/AgCl layer (e.g. byapplying Ag/AgCl ink or by sputter-coating a Ag or Ag/AgCl layer) orother reference electrode materials instead of a redox mediator.

As mentioned earlier, oxidizable interferents such as ascorbic acid,uric acid and acetaminophen, to name a few, cause inaccurate readings inthe output of an electrochemical biosensor. The present inventionreduces this effect considerably by using an applied potential thatminimizes oxidaton of these interferents. Also important is thecomposition of the reagents disposed on W1 and W2. The reagents aredesigned to have a minimal effect on the response of the interferenceswhich also contributes to the accuracy of the analyte measurement.

All of the advantages of the present invention will be made clearer uponreview of the detailed description, drawings and appended claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of the four-layer embodiment of the presentinvention showing the open end, the vent and the electrical contactpoints of the laminated body.

FIG. 2 is an exploded, perspective view of the four-layer embodimentshowing the various layers of the laminated body.

FIG. 3 is a perspective view of the three-layer embodiment of thepresent invention showing the open end, the vent and the electricalcontact points of the laminated body.

FIG. 4 an exploded, perspective view of the three-layer embodimentshowing the various layers of the laminated body.

FIG. 5 is a cross-sectional view of the present invention of FIG. 3.

FIGS. 6A, 6B, 6C, and 6D are top views of a strip of each layer of thepresent invention showing the patterns for making multiple sensors ofthe four-layer embodiment.

FIG. 6E is a top view of a segment of the laminated strip of the presentinvention showing the patterns for making multiple sensors of thefour-layer embodiment.

FIGS. 7A and 7B displays response curves for a creatinine sensor of thepresent invention in phosphate buffer solution.

FIG. 8 is a response curve using creatinine sensors of the presentinvention for blood samples.

FIG. 9 is a response curve of a creatinine sensor of the presentinvention showing the response to creatine and creatinine.

FIG. 10 is a graph of the response to volume sample of a creatininesensor of the present invention.

FIG. 11 is a response curve of a glucose sensor of the present inventionin phosphate buffer.

FIG. 12 is a response curve of a glucose sensor of the present inventionin urine.

FIG. 13 is a response curve of a cholesterol sensor of the presentinvention in phosphate buffer.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

The preferred embodiment of the present invention is illustrated inFIGS. 1-13. FIG. 1 shows a sensor 10 of the present invention using the4-layer construction. Sensor 10 has a laminated body 100, a fluidsampling end 110, an electrical contact end 120, and a vent opening 52.Fluid sampling end 110 includes a sample fluid channel 112 between asampling end aperture 114 and vent opening 52. Electrical contact end120 has at least three discreet conductive contacts 122, 124 and 126.

Referring now to FIG. 2, laminated body 100 is composed of a baseinsulating layer 20, a first middle layer 30, a second middle layer 40,and a top layer 50. All layers are made of a dielectric material,preferably plastic. Examples of a preferred dielectric material arepolyvinyl chloride, polycarbonate, polysulfone, nylon, polyurethane,cellulose nitrate, cellulose propionate, cellulose acetate, celluloseacetate butyrate, polyester, acrylic and polystyrene. Base insulatinglayer 20 has a conductive layer 21 on which is delineated a firstconductive 22, a second conductive 24 and a third conductive 26.Conductive conduits 22, 24 and 26 may be formed by scribing or scoringthe conductive layer 21, as illustrated in FIG. 2, or by silk-screeningthe conductive conduits 22, 24 and 26 onto base layer 20. Scribing orscoring of conductive layer 21 may be done by mechanically scribing theconductive layer 21 sufficiently to create the three independentconductive conduits 22, 24 and 26. The preferred scribing or scoringmethod of the present invention is done by using a carbon dioxide (CO₂)laser, a YAG laser or an eximer laser. An additional scoring line 28(enlarged and not to scale; for illustrative purposes only) may be made,but is not necessary to the functionality of sensor 10, along the outerbe edge of base layer 20 in order to avoid potential static problemswhich could give rise to a noisy signal. Conductive layer 21 may be madeof any electrically conductive material like copper, gold, tinoxide/gold, palladium, other noble metals or their oxides, or carbonfilm compositions, preferably gold or tin oxide/gold. A useable materialfor base layer 20 is a tin oxide/gold polyester film (Cat. No. FM-1) ora gold polyester film (Cat. No. FM-2) sold by Courtaulds PerformanceFilms, Canoga Park, Calif.

First middle layer 30 has a first electrode cutout 32 which exposes aportion of first conductive 22, a second electrode cutout 34 whichexposes a portion of second conductive 24 and a third electrode cutout36 which expose a portion of third conductive 26. First middle layer 30is made of a plastic material, preferably a medical grade one-sided tapeavailable from Adhesive Research, Inc., of Glen Rock, Pa. Acceptablethicknesses of the tape for use in the present invention are in therange of about 0.001 in. (0.025 mm) to about 0.005 in. (0.13 mm). Onesuch tape, Arcare® 8666 (about 0.003 in. (0.075 mm)), was preferredbecause of its ease of handling and it showed good performance in termsof its ability to hold a sufficient quantity of chemical reagents and topromote capillary action through sample fluid channel 112 of sensor 10.It should be understood that the use of a tape is not required. Aplastic insulating layer may be coated with a pressure sensitiveadhesive, or may be ultrasonically-bonded to base layer 20, or may besilk-screened onto base layer 20 to achieve the same results as usingthe polyester tape mentioned.

The three cutouts 32, 34 and 36 define electrode areas W1, R and W2,respectively, and hold chemical reagents forming two working electrodesand one in reference electrode. For biosensors measuring analytes suchas glucose and cholesterol, only two cutouts are required that holdchemical reagents for a working electrode and a reference electrode.Typically, electrode area R must be loaded with a redox reagent ormediator to make the reference electrode function. If R is not loadedwith a redox reagent or mediator, working electrodes W1 and W2 will notwork properly. The redox reagent preferably contains an oxidized form ofa redox mediator, at least one binder, and a surfactant. R could also beloaded or coated with silver/silver chloride or other referenceelectrode materials.

Examples of useable redox mediators are Fe(CN)₆ ³⁻, Fe(CN)₆ ⁴⁻,Fe(phen)₃ ²⁺(phen=1,10-phenanthroline), Fe(bpy)₃ ²⁺(bpy=2,2′-bipyridine), Co(NH₃)₆ ²⁺, Co(phen)₃ ²⁺, Co(bpy)₃ ²⁺,Os(bpy)₂Cl⁺, Os(phen)₂Cl³⁰ Ru(bpy)₂ ²⁺, Rb(bpy)₂ ²⁺, cobaltphthalocyanine, various ferrocenes, methylene blue, methylene green,7,7,8,8-tetracyanoquinodimethane (TCNQ), tetrathiafulvalene (TTF),toluidine blue, meldola blue, N-methylphenazine methosulfate,phenyidiamines, 3,3′,5,5′-tetramethylbenzidine (TMB), pyrogallol, andbenzoquinone (BQ). Silver/silver chloride or other reference electrodematerials could also be used.

The redox mediator may be any inorganic or organic redox species. Themediator may also be in either the reduced or oxidized form. Because alow applied potential (−0.15 V) is used in the present invention fordetecting the reduction current signal of the product of the enzymaticreaction, if a mediator is used for the reference electrode instead ofAg/AgCl, a reduced form of the redox mediator is preferred for thereference electrode. Use of the reduced form of the redox mediator willminimize carry-over from the reference electrode R to the workingelectrodes, W1 and W2, which is more likely to occur if an oxidized formof the redox mediator is used at the reference electrode R.

It is preferable that the mediator is capable of being oxidizedchemically by hydrogen peroxide resulting from enzymatic reactions suchas those illustrated in Eqs. (1) to (3) or Eq. (4) above. It is furtherdesirable that the oxidized form of the mediator is capable of beingreduced electrochemically at the working electrodes at the appliedpotential. It is still further desirable that the mediator is stable inthe matrix. The preferred mediator in the present invention is potassiumferrocyanide (K₄Fe(CN)₆). The preferred binders are polyethylene oxideand various water soluble cellulose materials like methyl cellulose andthe preferred surfactant is t-octylphenoxypolyethoxyethanol.

Generally, electrode area W1 is loaded with a reagent containingchemical components similar to that loaded in electrode area R. Thesesimilarities will become clearer to those skilled in the art when thereagent mixes are later described in more detail. The difference betweenthe reagents loaded in W1 and R is that the reagent loaded in electrodearea W1 also contains a peroxidase capable of being catalyticallyreactive with the mediator and at least one enzyme capable of catalyzinga reaction involving the analyte to be measured.

For a creatine sensor, the reagent preferably contains three enzymes, areduced form of a redox mediator, at least one binder, and a surfactant.The enzymes are preferably creatine amidinohydrolase (C2) sarcosineoxidase (SO) and the peroxidase. The peroxidase may be from any sourcesuch as soybean (soybean peroxidase (SBP)) or horseradish root(horseradish root peroxidase (HRP)).

For a glucose sensor, the reagent preferably contains two enzymes, areduced form of a redox mediator, at least one binder, and a surfactant.The enzymes are preferably glucose oxidase (GOD) and the peroxidasementioned above.

For a cholesterol sensor, the reagent preferably contains three enzymes,a reduced form of a redox mediator, at least one binder, and asurfactant. The enzymes are preferably cholesterol esterase, cholesteroloxidase and the peroxidase mentioned above.

For the creatinine sensor, electrode area W2 is preferably loaded withthe same chemical reagents loaded into electrode area W1 but with theaddition of another enzyme (fourth enzyme). This other enzyme is alsocapable of catalyzing a reaction involving a substrate for the enzyme.The mediator must be capable of transferring electrons transferredbetween the enzyme-catalyzed reaction and the working electrode tocreate a current representative of the concentration of the substrateand, more specifically, representative of the concentration ofcreatinine. The fourth enzyme is preferably creatinine amidohydrolase(C1).

The cutouts and electrode areas of first middle layer 30 are positionedrelative to each other and to the flow of the sample fluid in samplefluid channel 112 such that the possible carryover from one electrodearea to another electrode area could be minimized. Using fluid sampleend 110 of sensor 10 as a reference point, the arrangements of theelectrode areas could be W1-W2-R, W1-R-W2, R-W1-W2, W2-W1-R, W2-R-W1, orR-W2-W1. The preferred position for analytes having more than onecompeting substrate reaction such as the creatinine sensor was found tobe W1-R-W2. The preferred position for analytes having only onesubstrate reaction such as glucose and cholesterol was found to beR-W1-W2.

Second middle layer 40 has a U-shaped channel cutout 42 located atsecond layer sensor end 41. The length of channel cutout 42 is such thatwhen second middle layer 40 is layered on top of first middle layer 30,electrode areas W1, W2 and R are within the space defined by channelcutout 42. The thickness of second middle layer 40 was found to beimportant for the speed of the sample fluid flow into sample fluidchannel 112, which is filled by capillary action of the sample fluid.

Top layer 50, which is placed over second middle layer 40, has a ventopening 52 spaced from fluid sample end 110 of sensor 10 to insure thatsample fluid in fluid channel 112 will completely cover electrode areasW1, W2 and R. Vent opening 52 is placed in top layer 50 so that i willalign somewhat with the bottom of channel cutout 42 of second middlelayer 40. Preferably, vent opening 52 will expose a portion of andpartially overlay the bottom of the U-shaped cutout 42 of second middlelayer 40.

FIG. 3 shows another embodiment of the present invention showing asensor 500 of the present invention using 3-layer construction. Sensor500 has a laminated body 600, a fluid sampling end 610, an electricalcontact end 620, and a vent opening 542. Fluid sampling end 610 includesa sample fluid channel 612 between a sampling end aperture 614 and ventopening 542. Electrical contact end 620 has three discreet conductivecontacts 622, 623 and 624.

Referring now to FIG. 4, laminated body 600 is composed of a baseinsulating layer 520, a middle layer 530, and a top layer 540. Alllayers are made of a dielectric material, preferably plastic. Baseinsulating layer 520 has a conductive layer 521 on which is delineated afirst conductive conduit 522, a second conductive conduit 523 and athird conductive conduit 524. Conductive conduits 522, 523 and 524 maybe formed by scribing or scoring the conductive layer 521 as illustratedin FIG. 4 and shown as scribe line 527 and 528 or by silk-screening theconductive conduits 522, 523 and 524 onto base layer 520. Scribing orscoring of conductive layer 521 may be done by mechanically scribing theconductive layer 521 sufficiently to create the three independentconductive conduits 522, 523 and 524. The preferred scribing or scoringmethod of the present invention has been previously disclosed. Anadditional scoring line 529 (enlarged and not to scale; for illustrativepurposes only) may be made, but is not necessary to the functionality ofsensor 500, along the outer edge of base layer 520 in order to avoidpotential static problems which could give rise to a noisy signal.

Middle layer 530 has a U-shaped channel cutout 532 located at middlelayer sensor end 531. The length of channel cutout 532 is such that whenmiddle layer 530 is layered on top of base layer 520, electrode areas W,R and W₀ are within the space defined by channel cutout 532. Thethickness of middle layer 530 was found to be important for the speed ofthe sample fluid flow into sample fluid channel 612, which is filled bycapillary action of the sample fluid. Channel cutout 532 along with thebase layer 520 holds the reagent matrix 550, more clearly shown in FIGS.3-5. Channel cutout 532 also defines the area of the working electrode,the reference electrode and the second electrode. Electrode areas W, W₀and R are loaded preferably with the same chemical reagent. The reagentspreferably contain a reduced form of a redox mediator, at least onebinder, a surfactant, and at least one enzyme. Top layer 540, which isplaced over and coextensive with middle layer 530, has a vent opening542 spaced from fluid sample end 610 of sensor 500 to insure that samplefluid in fluid channel 612 will completely cover electrode areas W, Rand W₀. Vent opening 542 is placed in top layer 540 so that it willalign somewhat with the bottom of channel cutout 532 of middle layer530, the bottom meaning the channel cutout 532 located furthest fromsensor end 531. Preferably, vent opening 542 will expose a portion ofand partially overlay the bottom of the U-shaped cutout 532 of middlelayer 530. FIG. 5 shows an enlarged cross-sectional view of the variouslayers of the present invention. The layers are not to scale in orderthat the relationship of each component of the present invention may bebetter understood by those skilled in the art, especially scribe lines27 and 28. The possible electrode arrangements within the sample fluidchannel may be W-R-W₀, W-W₀-R, R-W-W₀, R-W₀-W, W₀-W-R or W,-R-W with thearrangement listed as the arrangement of electrodes would appear fromthe open end of the laminated body to the vent opening. The preferredposition was found to be R-W-W₀; that is, as the sample fluid enteredthe open end of the laminated body, the fluid would cover R first, thenW, then W₀.

The second electrode, W₀, is positioned so that the sample fluid reachesft last. The resulting current at W₀ thus triggers the reading meter tostart the measurement and analyte concentration determination process.Such an arrangement obviates reliability and accuracy problems due to aninsufficient sample fluid size. It should be pointed out that W₀ canalso be used as a counter electrode. The resulting three-electrodesystem (i.e. working electrode, reference electrode and counterelectrode) would be used in the case of a sample fluid having highresistance. It should also be pointed out that W₀, combined with R, canbe used to measure the resistance of the sample fluid. The resultingresistance could be used to estimate the hematocrit of a blood sampleand therefore to correct the measurement for hematocrit effect.

Creatinine Sensor Preparation of Reagents 1, 2 and 3

Reagents 1, 2 and 3 comprise the reduced form of a redox mediator, abinder, and a surfactant. The reduced form of the redox mediator must bestable in the reagent matrices and must make the reference electrodefunction well. Its quantity in the formulation must be sufficient toattain a working linear range. The preferred redox mediator is potassiumferrocyanide. The binder should be sufficiently water-soluble and shouldalso be capable of stabilizing and binding all other chemicals in thereagents in electrode areas W1, W2 and R to the conductive surface/layer21 of base layer 20. The binders are polyethylene oxide and variouswater soluble cellulose materials. The preferred binder is methylcellulose and is available as Methocel 60 HG (Cat. No. 64655, FlukaChemical, Milwaukee, Wis.). Preferably, a small amount of anti-oxidantis added to Reagents 1, 2 and 3. The anti-oxidant stabilizes the redoxmediator, thus providing for a long-term shelf-life. The anti-oxidantmust not interfere with the enzymatic reactions (Eqs. (1) to (4)) andthe ensuing amperometric measurement. The preferred anti-oxidant issodium sulfite and is available from most chemical supply companies. Thesurfactant is necessary to facilitate dispensing of Reagents 1, 2 and 3into the cutouts for W1, W2 and R as well as for quickly dissolving thedry chemical reagents. The amount and type of surfactant is selected toassure the previously mentioned function and to avoid a denaturingeffect on the enzymes. The preferred surfactant is a polyoxyethyleneether. More preferably, it is t-octylphenoxypolyethoxyethanol and isavailable under the brand name Triton X-100.

Reagent 2, in addition to the components in Reagent 1, contains creatineamidinohydrolase (C-IIAT, 14 U/mg, Kikkoman, Japan), sarcosine oxidase(SOD-TE, about 33 U/mg, Kikkoman, Japan) and soybean peroxidase (SBP-MD,about 220 U/mg, Organic Technologies, Columbus, Ohio). Reagent 3, inaddition to the components in Reagent 2, contains creatinineamidohydrolase (C-IE, about 600 U/mg, Kikkoman, Japan). The reagents areprepared as follows:

Reagent 1

-   Step 1: Prepare a 1% (W/W) Methocel 60 Hg solution by stirring 1    gram of Methocel 60 HG in 100 ml of water for 4 hours.-   Step 2: Add 0.2 ml of 10% Triton X-100 into the methocel solution    from Step 1.-   Step 3: While stirring, add 2 grams of potassium ferrocyanide and    0.05 gram sodium sulfite to the solution from Step 2.

Reagent 2

-   Step 1-Step 3: Same as Reagent 1.-   Step 4: While stirring, add 0.5 gram of soybean peroxidase to the    solution from Step 3.-   Step 5: Add 2 gram of creatine amidinohydrolase to the solution from    Step 4.-   Step 6: Add 0.5 gram of sarcosine oxidase to the solution from Step    5.

Reagent 3

-   Step 1-Step 6: Same as Reagent 2.-   Step 7: While stirring, add 0.4 gram of creatinine amidohydrolase to    the solution from Step 6.

Creatinine Electrode Construction

A piece of a gold or tin oxide/gold polyester film available fromCourtaulds Performance Films is cut to shape, as illustrated in FIG. 2,forming base layer 20 of sensor 10. A conductive side of the gold or tinoxide/gold polyester film is scored. Scribing or scoring the conductivelayer may be done mechanically, by laser or by any other method tocreate three independent conductive paths. Preferably, a YAG, eximer orCO₂ laser is used. More preferably, the conductive layer is scored byCO₂ laser (25W laser available from Synrad, Inc., San Diego, Calif.). Asillustrated in FIG. 2, the film is scored by the laser such that threeelectrodes at sample fluid end 110 and three contact points 122, 124 and126 are formed at electrical contact end 120. The scoring line is verythin but sufficient to create three separate electrical conductors. Ascoring line 28 can be made, but is not necessary, along the outer edgeof base layer 20 to avoid potential static problems which could cause anoisy signal from the finished sensor 10.

As mentioned earlier, the conductive conduits may be deposited on theinsulating layer by screen printing, by vapor deposition, or by anymethod that provides a conductive layer which adheres to the baseinsulating layer. Other conductive coatings may also be used such aspalladium film or other noble metal film or their oxides or a carbonfilm composition.

A piece of one-sided adhesive tape, having a thickness preferably ofabout 0.0025 in. (0.06 mm), is then cut to size and shape forming firstmiddle layer 30 so that it will cover a majority of the conductive layer21 of base layer 20 except for exposing a small electrical contact areaillustrated in FIG. 1. Three rectangular, square or circular cutouts 32,34 and 36 of substantially equal size are punched by CO₂ laser ordie-cut. Cutouts 32, 34 and 36 define the electrode areas W1, W2 and R,which hold chemical reagents. The size of the cutouts is preferred to bemade as small as possible in order to make the fluid sample channel 112of sensor 10 as short as possible while still being capable of holdingsufficient chemical reagent to function properly. The preferred hole isround in shape and has a diameter of about 0.043 in. (1.1 mm). Asillustrated in FIG. 2, cutouts 32, 34 and 36 are aligned with each otherand have a spacing of about 0.026 in. (0.65 mm) between them. Thecircular cutouts are for illustrative purposes only. It should beunderstood that the shape of the cutouts is not critical provided thatthe size of the cutouts is big enough to facilitate dispensing chemicalreagents but small enough to allow for a reasonably small samplechannel. As stated previously, the preferred arrangement of theelectrodes formed in cutouts 32, 34 and 36 is W1 (working electrode 1),R (reference electrode) and W2 (working electrode 2). The surface of thefirst middle layer 30 must be sufficiently hydrophilic. This is achievedby coating a layer of hydrophilic polymer or surfactant onto the firstmiddle layer. Preferably, the non-sticky side, which will face up to thefluid channel, is treated or coated with a surfactant (e.g. 0.05 %Triton X-100). A piece of the one-sided adhesive tape is cut to shape asdepicted in FIG. 1 so that it will cover a majority of conductive baselayer 20 except for exposing a small electric contact area (3×6 mm) andthree cutouts defining electrode areas R, W1 and W2.

0.5 microliter of Reagent 1 is dispensed into electrode area R. Reagent1 is a mixture of a redox mediator, a binder/stabilizer and asurfactant. The preferred mixture for Reagent 1 is made by mixing thefollowing components in the described percentages (W/W%): about 2%potassium ferrocyanide, about 1% Methocel 60 HG, about 0.05% sodiumsulfite, about 0.02% Triton X-100.

0.5 microliter of Reagent 2 is dispensed into electrode W1. Reagent 2 isa mixture similar to that of Reagent 1 but with the addition of threeenzymes, i.e. C2, SO and SBP, capable of catalyzing a reaction involvinga substrate of creatine. The preferred mixture for Reagent 2 is made bymixing the following percentages (W/W%) of the following ingredients:about 2% C2, about 0.5% SO, about 0.5% SBP, about 2% potassiumferrocyanide, about 1% Methocel 60 HG, about 0.05% sodium sulfite, about0.02% Triton X-100.

0.5microliter of Reagent 3 is dispensed into electrode W2. Reagent 3 isa mixture similar to that of Reagent 2 but with the addition of anenzyme, e.g. C1, capable of catalyzing a reaction involving a substrateof creatinine. The preferred mixture for Reagent 3 is made by mixing thefollowing percentages (W/W%) of the following ingredients: about 0.4%C1, about 2% C2, about 0.5% SO, about 0.5% SBP, about 2% potassiumferrocyanide, about 1% Methocel 60 HG, about 0.05% sodium sulfite, about0.02% Triton X-100.

After the addition of the reagents to the electrode areas, the device isdried in an oven for about 5 minutes at 37° C. After drying, a piece ofdouble-sided tape, having a thickness preferably of about 0.007 in.(0.18 mm) and available from Adhesive Research, Inc. (Cat. No. X12314),is cut into shape with a U-shape notch cutout at one end as illustratedin FIG. 2. The preferred size of the cutout is about 0.264 in. (6.7 mm)long by about 0.065 in. (1.65 mm) wide. The double-sided tape serves asa spacer layer and is second middle layer 40. The U-shape cutout is madewith the CO₂ laser described earlier. The thickness of the double-sidedtape along with the length and width of the cutout defines the volumesize of the fluid sample channel and the relative speed the fluid samplemoves into the defined chamber. The preferred size of the U-shapedcutout 42 is about 0.264 in. long (6.7 mm), about 0.065 in. wide (1.65mm) and about 0.007 in. thick (0.18 mm).

A piece of a transparency film having a thickness preferably of about0.0043 in. (0.11 mm)(Cat. No. PP2200 or PP2500 available from 3M) isfashioned into top layer 50. A rectangular vent hole 52 is made usingthe CO₂ laser or die-cut previously mentioned. The preferred size ofvent hole 42 is about 0.065 in. (1.65 mm) by about 0.059 in. (1.50 mm).The center of vent hole 52 is located approximately 0.234 in. (5.95 mm)from fluid end 110 of sensor 10. Top layer 50 is aligned and layeredonto second middle layer 40 to complete the assembly, as illustrated inFIG. 1, of sensor 10. Although the preferred embodiment is described, itshould be understood that the present invention may have a variety ofembodiments without detracting from the spirit of the present invention.

When a fluid sample is applied to a single strip of the presentinvention, the fluid sample enters the channel through sampling endaperture 114 and flows over W1, R and W2 and stops at the threshold ofvent opening 52. The length of the fluid channel 112, i.e. from samplingend aperture 114 to the threshold of vent opening 52, is about 0.208 in.(5.2 mm). The volume of the channel is calculated to be 1.54microliters.

Although the description of electrode construction above describesconstruction for a single sensor, the design and materials used areideal for making multiple sensors from one piece of each layer materialas shown in FIGS. 6A-6E. This would be accomplished by starting with arelative large piece of base layer 20 having conducting layer 21thereon. A plurality of scored lines are made into conductive layer 21such that a repetitive pattern, as illustrated in FIG. 6A, is createdusing the preferred scribing method described previously whereby eachpattern will eventually define the three conductive paths 22, 24 and 26for each sensor. Similarly, a large piece of first middle layer 30,which is illustrated in FIG. 6B and which also has a plurality ofcutouts 32, 34, and 36 in a repetitive pattern, is sized to fit overbase layer 20 in such a way that a plurality of sensors 10 will be hadwhen completed. The size of each cutout and the electrode materialdisposed in the plurality of electrode areas W1, R and W2 are similar tothat disclosed above.

After disposing Reagents 1, 2 & 3 in their respective cutouts and dried,a large piece of second middle layer 40 having a plurality of elongatedcutouts 42 and illustrated in FIG. 6C is layered onto first middle layer30 such that each elongated cutout 42 of second middle layer 40 containscorresponding cutouts 32, 34 and 36 of first middle layer 30. Acomparably-sized top layer 50 having a plurality of vent openings 52 ina repetitive pattern, as shown in FIG. 6D, is layered onto second middlelayer 40. FIG. 6E is a top view of the combined layers. The laminatedstrip created by the four layers 20, 30, 40 and 50 has a plurality ofsensors 10 that can be cut from the laminated strip. The laminated stripis cut longitudinally along line A—A′ at fluid sampling end 210 to forma plurality of sampling apertures 114 and longitudinally along line B—B′at electrical contact end 220 to form a plurality of conductive contacts122, 124 and 126. The laminated strip is also cut at predeterminedintervals along line C—C′ forming a plurality of individual sensors 10.

Shaping of the fluid sampling end 120 of each sensor 10, as illustratedin FIG. 1, may be performed if desired. It should be understood by thoseskilled in the art that the order in which the laminated strip can becut is not important. For instance, the laminated strip may be cut atthe predetermined intervals (C—C′) and then the cuts along A—A′ and B—B′can be made to complete the process.

Chronoamperometry (I-t curve) was used for measurement of the currentresponse of the strips using an Electrochemical Analyzer (CHInstruments, Model 812, Austin, Tex.). If not stated otherwise, thecurrent at 20 seconds was recorded.

The following examples illustrate the unique features of the presentinvention. A potential of −0.15 Volts was applied across the workingelectrodes and the reference electrode. The resultant current signalsbased on the oxidation of the reduced form of the redox mediator arerepresentative of the creatine and creatinine concentrations inaccordance with the preferred embodiment of the present invention.

EXAMPLE 1 Demonstration of Linear Range

Sample strips of the present invention were first tested in phosphatebuffer solution (PBS) containing 0 to 5 mg/dL creatinine with anElectrochemical Analyzer (CH Instruments, Model 812, Austin, Tex.).Table 1A shows the measured current response in nanoamperes of a sensorof the present invention to varying concentrations of creatinine inphosphate buffer solution.

TABLE 1A Response for Creatinine in PBS Concentration (mg/dL) Current(nA) 0.0 0.0 0.2 4 0.4 12.5 0.6 25 1 51 2 123 3 182 5 263 10 472 15 520

A graphical representation of the above data is shown in FIGS. 7A and7B. FIG. 7A is an enlarged view of graphical representation forcreatinine concentrations of 0 to 1 mg/dL. As seen from the data and thegraphs, the sensors of the present invention respond to small amounts ofcreatinine as low as 0.2 mg/dL. The sensors also exhibit a linearrelationship of current response versus creatinine concentration over aconcentration range from about 0.2 to about 10 mg/dL.

In order to test the response of the strips in a real sample, a sampleof venous blood was collected and separated into several aliquots. Eachaliquot was spiked with different creatinine concentrations ranging from0 to 25 mg/dL. The aliquots were each measured using a sensor of thepresent invention with the Electrochemical Analyzer. Table 1B shows thecurrent response in nanoamps in a blood sample spiked with varyinglevels of creatinine.

TABLE 1B Response for Creatinine in Blood mg/dL (C) nA (1) 0 21 1 46 272 5 150 40 310 15 440 20 536 25 623

A graphical representation of the test data is shown in FIG. 8. The testresults indicate that the sensors of the present invention have a linearresponse (current response vs. creatinine concentration) over acreatinine concentration range from about 0 to about 20 mg/dL, butcontinue to respond above this range.

EXAMPLE 2 Demonstration of Precision of Sensors

the precision of the sensors of the present invention was investigatedat the creatinine level of 2 mg/dL in a blood sample. The creatininelevel was confirmed by measurement using the Nova Stat Profile M, NovaBiomedical Co., Waltham, Mass. Table 2 shows the current response innanoamps of a blood sample spiked with 2 mg/dL of creatinine usingvarious sensors of the present invention.

TABLE 2 Precision Study Run No. Current (nA) 1 55 2 55 3 55 4 61 5 55 655 7 59 8 53 9 59 10 56 11 56 12 59 13 60 14 62 15 62 16 62 17 60 18 5719 60 20 62 21 59 Average   58.2 CV, %   5.0

Twenty-one (21) sensor strips from the same batch were tested and acoefficient of variation (CV) was found to be 5.0%.

EXAMPLE 3 Demonstration of Interference Free Feature

The most important challenge for the measurement of creatinine is theinterference from creatine, as it always co-exists in the sample alongwith creatinine. The unique design of the present invention makes itpossible to eliminate the Interference from creatine. This is achievedby subtracting the response obtained at W1 from the response obtained atW2, and is represented by the following equation:I=Iw₂−Iw₁  Eq. (7)where Iw₂ is the current at W2 (second working electrode)

-   -   Iw₁ is the current at W1 (first working electrode)    -   I is the difference between W2 and W1 and represents the current        due to oxidation of the mediator of its reduced form, which is        proportional to the creatinine concentration in the sample

Because W1 and W2 have the same surface area, the potential interferencepresent in the sample fluid should give relatively identical signalsfrom each working electrode.

This was tested by spiking a blood sample with different concentrationsof creatine, i.e. (A) blood sample (1.0 mg/dL creatinine, measured withNova Stat Profile M, Nova Biomedical Corporation, Waltham, Mass.); (B)same as (A) but with addition of 5 mg/dL creatine; (C) same as (A) butwith the addition of 10 mg/dL creatine. FIG. 9 displays the effect ofadded creatine on the current response of the strips of the presentinvention. It is noticed that, although the sensor output currentsincrease upon spiking with creatine, the current difference at W2 andW1, representative of the analyte creatinine, remains nearly unchanged.

Other common interferences are from oxidizable substances such asascorbic acid and acetaminophen present in the sample. A blood samplewas spiked with different levels of ascorbic acid and acetaminophen.Table 3 shows the test results obtained.

TABLE 3 Response Change Upon Addition of Interferent mM Current (nA) 2mg/dL creatinine: ascorbic acid 0.0 26 0.1 25 0.2 21 2 mg/dL creatinine:acetaminophen 0.0 32.5 1.0 34

The result shows that less than 0.1 mM ascorbic acid and 1 mMacetaminophen will not influence the measurement of creatinine due tothe low level of applied potential (−0.15 V) previously described. Thesesubstances are not oxidized at that level of applied potential.

EXAMPLE 4 Demonstration of Minimum Sample Volume Feature

The unique design of the present invention enables the measurement ofsample sizes smaller than have heretofore been possible. Blood samplesare applied to the sensors and the samples travel along the sample fluidchannel to the threshold of the vent hole. In order to test the volumeeffect on sensor response, different blood sample volumes were appliedto the sensors. Table 4 shows the current response versus volume size.

TABLE 4 Response to Sample Volume Volume (uL) Current (nA) 1.5 60 2 59 347 4 55 5 55

The resulting current signals were plotted against volume and is shownin FIG. 10. From the data and the graphical representation, sensors ofthe present invention for the sizes disclosed earlier show no dependenceof the response on the sample volume if the volume is above1.5microliters.

Glucose and Cholesterol Sensors Preparation of Reagents

Reagents for both Glucose and Cholesterol Sensors comprise the reducedform of a redox mediator, a peroxidase, at least one binder, asurfactant, and at least one analyte-based enzyme. The preferred redoxmediator is potassium ferrocyanide. The preferred peroxidase is soybeanperoxidase and is available as SBP-MD (about 220 U/mg, OrganicTechnologies, Columbus, Ohio). The preferred binder for the glucosesensor is methyl cellulose and is available as Methocel 60 HG (Cat. No.64655, Fluka Chemical, Milwaukee, Wis.). The preferred binder for thecholesterol sensor is also a cellulose material and is available asKlucel®-EF (Hercules, Wilmington, Del.). Preferably, a small amount ofanti-oxidant is added to the Reagents for the glucose and cholesterolsensors. The preferred anti-oxidant is sodium sulfite. The preferredsurfactant is a polyoxyethylene ether. More preferable, it ist-octylphenoxypolyethoxyethanol and is available under the brand nameTriton X-100.

For the glucose sensor, the analyte-based enzyme is glucose oxidase andis available as GO3AC from Biozyme, San Diego, Calif. For thecholesterol sensor, the analyte-based enzyme is a mix of cholesterolesterase available as COE-311 from Toyobo, Japan, and cholesteroloxidase available as COO-311, Toyobo, Japan.

The preferred reagent mixture for the glucose sensor is made by mixingthe following components in the described percentages (W/W%): About 0.5%glucose oxidase, about 0.5% soybean peroxidase, about 2% potassiumferrocyanide, about 1% Methocel 60 HG, about 0.1% sodium sulfite, andabout 0.02% Triton X-100.

The preferred reagent mixture for the cholesterol sensor is made bymixing the following components in the described percentages (W/W%):About 1% cholesterol esterase, about 2% cholesterol oxidase, about 0.5%soybean peroxidase, about 5% potassium ferrocyanide, about 1% Klucel-EF,about 0.1% sodium sulfite, and about 0.02% Triton X-100.

Glucose and Cholesterol Electrode Construction

The construction of the glucose and cholesterol sensors is based on thesecond embodiment previously described and illustrated in FIGS. 3 and 4.A piece of a gold or tin oxide/gold polyester film available fromCourtaulds Performance Films is cut to shape, as illustrated in FIGS. 3and 4, forming base layer 520 of sensor 500. A CO₂ laser is used toscore the gold or tin oxide/gold polyester film (25W laser availablefrom Synrad, Inc., San Diego, Calif.). As illustrated in FIG. 4, thefilm is scored by the laser creating scoring line 527 and 528 such thatthree electrodes at sample fluid end 610 and three contact points 622,623 and 624 were formed at electrical contact end 620. The scoring lineis very thin but sufficient to create two separate electricalconductors. An additional scoring line 529 made be made, but is notnecessary, along the outer edge of base layer 520 to avoid potentialstatic problems which could cause a noisy signal from the finishedsensor 500.

A piece of double-sided tape (Arcare® 7840) available from AdhesiveResearch, Glen Rock, Pa., is cut to size and shape forming middle layer530 with U-shaped channel 532 so that it will cover a majority of theconductive layer 521 of base layer 520 except for exposing a smallelectrical contact area at electrical contact end 620 illustrated InFIG. 3. The U-shaped channel 532 is cut by using the CO₂ laser. Middlelayer 530 is then layered onto base layer 520. As mentioned earlier,this middle layer 530 serves as a spacer and defines the size of thefluid sample channel 612. It also defines the electrode area 526 thatholds the electrode reagent matrix 550. Its width and length isoptimized to provide for a relatively quick moving fluid sample. Thepreferred size of U-shaped channel 532 is about 0.039 in. (1.0 mm) wideby about 0.134 in. (3.4 mm) long.

1.0 microliter of reagent mix is dispensed into channel 532 to formelectrode W, R and W₀. The reagent mix is a mixture of a redox mediator,a peroxidase, a binder, a surfactant, and at least one analyte-basedenzyme. The preferred composition for the reagent mix is made by mixingthe ingredients disclosed above for the glucose and cholesterol sensors.After the addition of the reagent mix, the devices were dried in oven at37° C. for about 5 minutes.

After drying, apiece of a transparency film (Cat. No. PP2200 or PP2500available from 3M) is fashioned into top layer 540. A rectangular venthole 542 is made using the CO₂ laser previously mentioned. The preferredsize of vent hole 542 is about 0.039 in. (1.0 mm) by about 0.051 in.(1.30 mm). Vent hole 542 is located approximately 0.087 in. (2.2 m) fromfluid end 610 of sensor 500. Top layer 540 is aligned and layered ontomiddle layer 530 to complete the assembly, as illustrated in FIG. 3, ofsensor 500.

EXAMPLE 5 Demonstration of Response to Glucose

Sample strips of the present invention were first tested in phosphatebuffer solution (PBS) containing 0 to 20 mg/dL glucose with anElectrochemical Analyzer (CH Instruments, Model 812, Austin, Tex.).Table 5A shows the measured current response in nanoamperes of a sensorof the present invention to varying concentrations of glucose inphosphate buffer solution.

TABLE 5A Response for Glucose in PBS Concentration (mg/dL) Current (nA)0.0 3 0.5 6 1 10 2 20 5 41 10 83 20 128

A graphical representation of the above data is shown in FIG. 11. Asseen from the data and the graph, the sensors of the present inventionrespond to small amounts of glucose and exhibit a near-linearrelationship of current response versus glucose concentration over aconcentration range from about 0.0 to about 20 mg/dL.

In order to test the response of the strips in a real sample, urine wascollected and separated into several aliquots. Each aliquot was spikedwith different glucose concentrations ranging from 0 to 50 mg/dL. Thealiquots were each measured using a sensor of the present invention withthe Electrochemical Analyzer. Table 5B shows the current response innanoamps in a urine sample spiked with varying levels of glucose.

TABLE 5B Response for Glucose in Urine mg/dL (C) nA (1) unspiked 10 1024 20 35 50 74 100  96

A graphical representation of the test data is shown in FIG. 12. Thetest results indicate that the sensors of the present invention have alinear response (current response vs. glucose concentration) over aglucose concentration range from about 0 to about 50 mg/dL, but continueto respond above this range.

EXAMPLE 6 Demonstration of Response to Cholesterol

Sample strips of the present invention were first tested in a Sigmacalibration standard diluted with phosphate buffer solution (PBS)containing 0 to 200 mg/dL glucose with an Electrochemical Analyzer (CHInstruments, Model 812, Austin, Tex.). Table 6 shows the measuredcurrent response in nanoamperes of a sensor of the present invention tovarying concentrations of cholesterol in phosphate buffer solution.

TABLE 6 Response for Cholesterol in PBS Concentration (mg/dL) Current(nA) 0 5 25 150 50 225 100 300 200 330

A graphical representation of the above data is shown in FIG. 13. Asseen from the data and the graph, the sensors of the present inventionrespond to small amounts of cholesterol and exhibit a near-linearrelationships of current response versus cholesterol concentration overa concentration range from about 0.0 to about 200 mg/dL.

1. A disposable electrode strip for measuring an analyte in a fluidsample said strip comprising: a laminated strip having a first stripend, a second strip and and a vent opening spaced from said first stripand, said laminated strip comprising a base layer with at least twoelectrodes delineated thereon, a reagent holding layer carried on saidbase layer, said reagent holding layer having at least two cutouts, achannel forming layer carried on said reagent holding layer, and acover; an enclosed channel between said first strip end and said ventopening, said enclosed channel containing said at least two cutouts; afirst reagent disposed in a first cutout of said at least two cutoutsforming a reference electrode, said first reagent comprising a referenceelectrode material selected from the group consisting of silver chloridewhen said reference electrode is silver and a mixture made by combiningat least a redox mediator and at least one binder when said referenceelectrode is selected from the group consisting of gold, gold/tin oxide,palladium, platinum and carbon composition; a second reagent disposed ina second cutout of said at least two cutouts forming a first workingelectrode, said second reagent comprising a redox mediator, at least onebinder, at least one enzyme that is a substrate of said analyte and aperoxidase capable of catalyzing a reaction involving said redoxmediator wherein said redox mediator is oxidized wherein said secondreagent is water soluble; and conductive contacts at said second stripend and insulated from said enclosed channel.
 2. The electrode strip ofclaim 1 further comprising a third cutout and a third reagent disposedin said third cutout forming a second working electrode wherein saidthird reagent comprises said redox mediator and said at least one binderwherein said third reagent is water soluble.
 3. The electrode strip ofclaim 2 wherein said third reagent further includes said at least oneenzyme, a substrate of said at least one enzyme and a peroxidase.
 4. Theelectrode strip of claim 2 wherein said first reagent, said secondreagent and said third reagent are made from a mixture having startingcomponents comprising about 1 wt % to about 6.5 wt % of said redoxmediator, about 1 wt % of binder, and about 0.02 wt % of said surfactantin water.
 5. The electrode strip of claim 4 wherein said first reagent,said second reagent and said third reagent further includes about 0.05wt % to about 0.1 wt % of an antioxidant.
 6. The electrode strip ofclaim 4 wherein said second reagent is made from a mixture havingstarting components in water comprising about 2 wt % of potassiumferrocyanide, about 1 wt % of methyl cellulose, about 0.02 wt % of saidt-octylphenoxypolyethoxyethanol, about 0.5 wt % of glucose oxidase, andabout 0.5 wt % of soybean peroxidase.
 7. The electrode strip of claim 4wherein said second reagent is made from a mixture having startingcomponents in water comprising about 5 wt % of potassium ferrocyanide,about 1 wt % of methyl cellulose, about 0.02 wt % oft-octylphenoxypolyethoxyethanol, about 2 wt % of cholesterol oxidase,about 1 wt % of cholesterol esterase, and about 0.5 wt % of soybeanperoxidase.
 8. The electrode strip of claim 4 wherein said secondreagent is made from a mixture having starting components in watercomprising about 2 wt % of potassium ferrocyanide, about 1 wt % ofmethyl cellulose, about 0.02 of t-octylphenoxypolyethoxyethanol, about 2wt % of creatine amidinohydrolase, about 0.5 wt % of sarcosine oxidase,and about 0.5 of soybean peroxidase, and wherein said third reagent ismade from a mixture having starting components in water comprising about2 wt % of said potassium ferrocyanide, about 1 wt % of said methylcellulose, about 0.02 wt % of said t-octylphenoxypolethoxyethanol, about2 wt % of said creatine amidinohydrolase, about 0.4 wt % of creatinineamidohydrolase, about 0.5 wt % of said sarcosine oxidase, and about 0.5wt % of said soybean peroxidase.
 9. The electrode strip of claim 8wherein said first reagent and said second reagent further includesabout 0.05 wt % of an antioxidant.
 10. The electrode strip of claim 2wherein the surface area of said first working electrode issubstantially the same size as the surface area of said second workingelectrode.
 11. The electrode strip of claim 1 wherein said peroxidase isat least one of soybean peroxidase and horseradish root peroxidase. 12.The electrode strip of claim 1 wherein said at least one enzyme is oneof creatine amidinohydrolase, glucose oxidase and cholesterol oxidase.13. The electrode strip of claim 12 wherein said second reagent furtherincludes a second enzyme when said at least one enzyme is one ofcreatine amidinohydrolase and cholesterol oxidase.
 14. The electrodestrip of claim 13 wherein said second enzyme is sarcosine oxidase whensaid at least one enzyme is creatine amidinohydrolase.
 15. The electrodestrip of claim 13 wherein said second enzyme is cholesterol esterasewhen said at least one enzyme is cholesterol oxidase.
 16. The electrodestrip of claim 1 wherein said redox mediator is an inorganic or organicredox species.
 17. The electrode strip of claim 16 wherein said redoxspecies is at least one of Fe(CN)₆ ³⁻, Fe(CN)₆ ⁴⁻,Fe(0,10-phenanthroline)₃ ²⁺, Fe(2,2′-bipyridine)₃ ²⁺, Co(NH₃)₆ ²⁺,Co(1,10-phenanthroline)₃ ²⁺, Co(2,2′-bipyridine)₃ ²⁺,Os(2,2′-bipyridine)₂Cl⁺, Os(1,10-phenanthroline)₂Cl⁺,Ru(2,2′-bipyridine)₂ ²⁺, Rh(2,2′-bipyridine)₂ ²⁺, cobalt phthalocyanine,ferrocenes, methylene blue, methylene green,7,7,8,8-tetracyanoquinodimethane, tetrathiafulvalene, toluidine blue,meldola blue, N-methylphenazine methosulfate, phenyldiamines,3,3′,5,5′-tetramethylbenzidine, pyrogallol, and benzoquinone.
 18. Theelectrode strip of claim 17 wherein said redox mediator is potassiumferrocyanide.
 19. The electrode strip of claim 1 wherein said enclosedchannel is hydrophilic.
 20. The electrode strip of claim 1 wherein saidenclosed channel has a volume of about 1.5 microliters.
 21. Theelectrode strip of claim 1 wherein said cover has a hydrophilic coatingon at least one side.
 22. The electrode strip of claim 1 wherein saidfirst reagent and said second reagent are made from a mixture havingstarting components in water comprising about 2 wt % of potassiumferrocyanide, about 1 wt % of methyl cellulose, about 0.02 wt % of said1-octylphenoxypolyethoxyethanol, about 0.5 wt % of glucose oxidase, andabout 0.5 wt % of soybean peroxidase.
 23. The electrode strip of claim22 wherein said first reagent and said second reagent further includesabout 0.1 wt % of an antioxidant.
 24. The electrode strip of claim 1wherein said first reagent and said second reagent are made from amixture having starting components in water comprising about 5 wt % ofpotassium ferrocyanide, about 1 wt % of methyl cellulose, about 0.02 wt% of t-octylphenoxypolyethoxyethanol, about 2 wt % of cholesteroloxidase, about 1 wt % of cholesterol esterase, and about 0.5 wt % ofsoybean peroxidase.
 25. The electrode strip of claim 24 wherein saidfirst reagent and said second reagent further includes about 0.1 wt % ofan antioxidant.
 26. A device for measuring creatinine comprising: acreatine measuring electrode having a creatine reagent comprising twoenzymes capable of catalyzing a reaction involving a substrate for theenzyme creatine, a redox mediator and a peroxidase enzyme capable ofcatalyzing a reaction involving said redox mediator wherein saidcreatine reagent is water soluble; a creatinine measuring electrodehaving a creatinine reagent comprising three enzymes capable ofcatalyzing a reaction involving a substrate for the enzyme creatinine, aredox mediator and a peroxidase enzyme capable of catalyzing a reactioninvolving said redox mediator wherein said creatinine reagent is watersoluble; and a reference electrode.
 27. The device of claim 26 whereinsaid two enzymes are creatine amidinohydrolase and sarcosine oxidase.28. The device of claim 26 wherein said three enzymes are creatinineamidohydrolase, creatine amidinohydrolase and sarcosine oxidase.
 29. Thedevice of claim 26 further comprising a sample chamber containing saidcreatine measuring electrode, said creatinine measuring electrode andsaid reference electrode.
 30. A creatinine sensor comprising: alaminated body having a fluid sampling end and an electrical contactend; a sampling end aperture at said fluid sampling end; a sample fluidchannel in communication between said fluid sample inlet and a ventopening, said sample fluid channel being adapted to collect a fluidsample through said sampling end aperture; a creatine measuringelectrode having a creatine reagent comprising two enzymes capable ofcatalyzing a reaction involving a substrate for the enzyme creatine, aredox mediator and a peroxidase enzyme capable of catalyzing a reactioninvolving said redox mediator wherein said creatine reagent iswater-soluble; a creatine-creatinine measuring electrode having acreatinine reagent comprising three enzymes capable of catalyzing areaction involving a substrate for the enzyme creatinine, a redoxmediator and a peroxidase enzyme capable of catalyzing a reactioninvolving said redox mediator wherein said creatinine reagent iswater-soluble; and a reference electrode wherein said creatine measuringelectrode, said creatine-creatinine measuring electrode and saidreference electrode are disposed within said sample fluid channel. 31.The creatinine sensor of claim 30 wherein said two enzymes are creatineamidinohydrolase and sarcosine oxidase.
 32. The creatinine sensor ofclaim 30 wherein said three enzymes are creatinine amidohydrolase,creatine amidinohydrolase and sarcosine oxidase.